The field of the invention is medical imaging methods and systems. More particularly, the invention relates to the imaging of a moving subject such as the beating heart.
When imaging a moving subject such as the beating heart, higher quality images can be obtained if the data is acquired very rapidly by the medical imaging system. With an x-ray CT system, for example, the x-ray source and detector are rotated around the subject to acquire a sufficient number of projection views from which a tomographic image can be reconstructed. This requires time and a trade-off is usually made between shortening the scan time by acquiring fewer projection views and image quality that improves with more projection views.
While this problem exists for x-ray CT, it is a much more significant problem in magnetic resonance imaging (MRI). When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “stipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A signal is emitted by the excited spins after the excitation signal B1 is terminated, this signal may be received and processed to form an image. When utilizing these signals to produce images, magnetic field gradients (Gx, Gy and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
The prevailing methods used to acquire NMR signals and reconstruct images use a variant of the well known Fourier transform (FT) imaging technique, which is frequently referred to as “spin-warp”. The spin-warp technique is discussed in an article entitled “Spin-Warp NMR Imaging and Applications To Human Whole-body Imaging” by W. A. Edelstein et al., Physics in Medicine and Biology, Vol. 25, pp. 751-756 (1980). It employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of NMR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (2DFT), for example, spatial information is encoded in one Cartesian coordinate system direction by applying a phase encoding gradient (Gy) along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient (Gx) in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse Gy is incremented (ΔGy) in the sequence of “views” that are acquired during the scan to produce a set of NMR data from which an entire image can be reconstructed.
To increase the rate at which image frames are acquired, image quality may be sacrificed by acquiring fewer phase encoding views, or by using faster pulse sequences that inherently result in lower quality images. With the spin-warp methods, therefore, there is a trade-off between the number of views that are acquired to achieve the desired image resolution and quality, and the rate at which NMR data for a complete image may be acquired.
More recently, an alternative method of acquiring NMR image data has been used in which no phase encoding gradients are employed. Instead, only a readout gradient is applied during the acquisition of each NMR signal (i.e., “view”) and a series of different views are acquired by rotating the angle of the readout gradient. Rather than sampling k-space in a rectilinear scan pattern as is done in Fourier imaging, this “projection reconstruction” method samples k-space with a series of views that sample radial lines extending outward from the center of k-space. The number of views needed to sample k-space determines the length of the scan and if an insufficient number of views are acquired, streak artifacts are produced in the reconstructed image.
Because the beating heart is constantly moving, the many different views needed to reconstruct an artifact-free image are acquired over a series of heart beats at approximately the same point, or “sphase”, in the cardiac cycle. Image acquisition is gated using an ECG trigger signal, and typically four to eight views (referred to as a “segment”) are acquired at a selected time interval after the cardiac trigger signal. The reconstructed image depicts the heart at a particular cardiac phase as determined by the selected delay time.
The standard backprojection method used in both MRI and x-ray CT is illustrated in FIG. 2. Each acquired signal projection profile 10 is backprojected onto the field of view 12 by projecting each signal sample 14 in the profile 10 through the FOV 12 along the projection path as indicted by arrows 16. In projecting each signal sample 14 in the FOV 12 we have no a priori knowledge of the subject and the assumption is made that the signals in the FOV 12 are homogeneous and that the signal sample 14 should be distributed equally in each pixel through which the projection path passes. For example, a projection path 18 is illustrated in FIG. 2 for a single signal sample 14 in one signal projection profile 10 as it passes through N pixels in the FOV 12. The signal value (P) of this signal sample 14 is divided up equally between these N pixels:Sn=(P×1)/N  (1)where: Sn is the NMR signal value distributed to the nth pixel in a projection path having N pixels.
Clearly, the assumption that the signal in the FOV 12 is homogeneous is not correct. However, as is well known in the art, if certain filtering corrections are made to each signal profile 10 and a sufficient number of filtered profiles are acquired at a corresponding number of projection angles, the errors caused by this faulty assumption are minimized and image artifacts are suppressed. In a typical, filtered backprojection method of image reconstruction, 400 projections are required for a 256×256 pixel 2D image and 203,000 projections are required for a 256×256×256 voxel 3D image. If the method described in the above-cited U.S. Pat. No. 6,487,435 is employed, the number of projection views needed for these same images can be reduced to 100 (2D) and 2000 (3D).
More than 20 years ago a method was proposed for reducing the number of projection views needed to produce adequate images of the beating heart, McKinnon and Bates “Towards Imaging The Beating Heart Usefully With A Conventional CT Scanner”, IEEE Transactions on Biomedical Engineering, Vol. BME-28, No. 2, Feb. 1981. The authors recognized that when acquiring views at different cardiac phases the stationary tissues surrounding the heart remained constant throughout and all the acquired views could be used to reconstruct a very high quality image of the stationary tissues. By combining the higher quality stationary tissue image data with the acquired moving tissue data an image could be reconstructed in which streak artifacts caused by stationary tissues could be removed. This method has not found significant clinical use, however, because the multi-source CT scanner for which the method was designed was not commercialized.
While a decent single-slice, 2D image may be acquired at one or more cardiac phases during a single breath-hold using known methods, prior methods are not fast enough to acquire a 3D image or multiple 2D slices at each cardiac phase during a single breath hold. Such images are necessary when the subject of the examination (such as coronary arteries) does not lie in a single 2D plane and either a multi-slice or 3D image acquisition is needed to make a diagnoses.